It is known that heating a cancer tumor to about 43.degree. C. can promote regression. Radiation, microwave and other methods of heating tissue have been used to promote tumor regression.
For example, dipole antennas have been inserted in catheters with electromagnetic energy radiated from antenna junctions. Typically such junctions are on the order of 1 mm. Energy radiated from individual antenna junctions may enhance or cancel each other depending on electromagnetic wave propagation characteristics within the tissue. Heating desired tumor locations is achieved through a power adjustment and sometimes phase control of the individual antennas. Because of small radiating antenna junctions, heating is usually limited to a few centimeters within the junction plane along the antennas. Temperature uniformity within such heated tumor volumes is dependent on wave propagation, tissue characteristics, thermal conductivity and blood perfusion rates, all of which are not directly controllable. Therefore this technique is usually limited to tumors of 3-4 cm long.
Use of radiofrequency induced currents with needles is another technique which has previously been attempted. Here, metallic needle pairs (electrodes) are implanted at approximately 1-2 cm spacings into a tumor volume and radiofrequency currents are passed through the electrode pairs. Heating is a function of tissue resistance to current flow between the electrodes which varies with different tissue types. Different techniques have been developed to drive currents through electrode pairs to achieve uniform therapeutic temperatures within tumor volumes. Since heating is a function of the electrical characteristics of tissue between the needle pairs, electrode spacing uniformity is critical. In practical applications where long electrodes are required, it is difficult to maintain necessary uniform spacings and hot spots can result which cause burns or related complications.
Another known technique utilizes stranded or solid wires of selected lengths, approximately 1-2 mm in diameter, inserted in arrays of catheters implanted within a tumor volume. The patient is then placed within an induction coil and exposed to high intensity magnetic fields. Different coil configurations have been used and driven in series or in parallel resonance circuit configurations to generate the required magnetic field. The implants are accordingly heated by resistive loses from any induced current circulations and the tumor tissue is heated by thermal conduction. Implant temperatures are achieved in accordance with Curie temperature characteristics of the ferromagnetic material used. The ferromagnetic property of these implants changes as a function of temperature, heating is gradually reduced as the Curie temperature is approached and further reduced when the Curie temperature is exceeded. Thermal regulation is dependent on a sharp transition in the Curie temperature curve at the desired temperature. The availability of implants that can be thermally regulated at desirable temperatures is limited by practical metallurgy limitations. Further, coils used to generate required high intensity magnetic fields are extremely inefficient. In fact, 1500-3000 Watts can be required and the implants need to be aligned with the applied magnetic field. Due to the high power requirements, both very expensive radiofrequency shielded rooms and complex cooling systems are required.
Yet another known system utilizes implantation of heating elements embedded within the walls of plastic catheters (typically 2.2 mm in diameter) which are then directly inserted into tissue. Heating is accomplished with an array of implanted catheters through thermal conduction. Temperature uniformity is maintained through active control of the current flowing in individual heating elements. Since direct current (dc) is used, the advantages of high frequency effects cannot be realized. Heating is only a function of the applied voltage and current. The technique is simple but there are practical limitations in the maximum voltage and current that can be applied to size limited embedded heating elements. Patient safety and electrical isolation requirements are more difficult to comply with for dc than with high frequency systems. Further, the resistance of heating elements is a function of length which when combined with voltage and current requirements restricts practical therapeutic applications.
Finally, catheters have been implanted in tumor tissue with heated water passed through the catheters to provide hyperthermia. Control of water temperature passing through individual catheters is used to maintain uniform tumor temperature. Clearly a very involved system and process. Further, there are unavoidable difficulties with miniaturized plumbing which limit practical applications. Changes in temperature within the small volume of water passing through the catheters due to thermal conduction results in undesirable temperature non-uniformity along the catheter lengths which additionally limits practical applications.